The present invention relates to the art of magnetic resonance imaging (MRI) of moving substances. It finds particular application in conjunction with imaging and measuring blood flow velocity by phase mapping and will be described with particular reference thereto. However, it is to be appreciated that the present invention is also applicable to imaging or phase mapping, velocity determination, acceleration determination, and the like of other anatomical movements, non-biological fluid flows, multi-component systems in which one gaseous, fluid, semi-fluid, solid, or other components moves relative to other components, or the like.
Magnetic resonance imaging provides a non-invasive means of imaging and measuring in clinical and other settings. Some of the many important applications offered by such technology are magnetic resonance flow measurement and magnetic resonance angiography.
Magnetic resonance (MR) flow quantitation scans are used to diagnose certain diseases by monitoring blood flow. See Bryant D. J., Payne J. A., Firmin D. N., and Longmore D. B., J. Comput Assist Tomogr, v. 8, p. 588 (1984); Underwood S. R., Firmin D. N, Klipstein R. H., Rees R. S. and Longmore D. B., Br Heart J., v. 57, p. 404 (1987); Klipstein R. H., Firmin D. N., Underwood S. R., Rees R. S. and Longmore D. B., Brit. Heart J., v. 58, p. 316 (1987). Similarly, magnetic resonance flow quantitation techniques are used to diagnose diseases by identifying cerebrospinal fluid flow patterns. See Quencer, R. M., et.al., Neuroradiology 32 (1990).
Heretofore, magnetic resonance imaging subjects have been positioned in a temporally constant magnetic field such that selected dipoles preferentially align with the magnetic field. A radio frequency pulse is applied to cause the preferentially aligned dipoles to resonate and emit magnetic resonance signals of a characteristic resonance radio frequency. The radio frequency magnetic resonance signals from the resonating dipoles are read out for reconstruction into an image representation.
In a Fourier transform imaging technique, a read gradient is applied during the read out of the echo for frequency encoding along a read axis and a phase-encode gradient is pulsed to step phase-encoding along a phase-encode axis between echoes. In this manner, each echo generates a data line in k-space. The relative phase-encoding of the data lines controls their relative position in k-space. Conventionally, the data line with zero phase-encoding extends across the center of k-space. Data lines with a phase-encoding gradient stepped in progressively positive steps are generally depicted as being above the center line of k-space; and, data lines with progressively negative phase-encoding steps are depicted as being below the center line of k-space. In this manner, a matrix, such as a 256.times.256 or a 512.times.512, etc., matrix of data values in k-space is generated. Fourier transformation of these values generates a conventional magnetic resonance image.
To strengthen the received magnetic resonance signals, the free induction decay magnetic resonance signal is commonly refocused into an echo. This may be done by reversing the polarity of a magnetic field gradient to induce a field or gradient echo. Analogously, the radio frequency excitation pulse may be followed with a 180.degree. radio frequency inversion pulse to refocus the signal as a spin echo. Moreover, by repeating the reversing of the magnetic field gradient, a series of gradient echoes can be generated following each radio frequency excitation pulse. Analogously, a series of spin echoes can be generated following each radio frequency excitation pulse by a series of the 180.degree. radio frequency refocusing pulses. As yet another option, a single radio frequency excitation pulse can be followed by a mixture of spin and gradient echoes. See, for example U.S. Pat. No. 4,833,408 of Holland, et al.
There are several known techniques for measuring flows using magnetic resonance experiments. For the purpose of measuring the flow velocity of blood or cerebrospinal fluid using magnetic resonance imaging, an image or a set of complex images are obtained with the flow information encoded in them. A wide variety of non-invasive methods are used for measuring flow. These methods can be summarized into three major categories: in-flow/out-flow; time-of-flight; and flow phase shift.
The in-flow/out-flow technique is based on the image intensity modulation caused by the excited spin moving out and un-excited spin moving in the imaging plane. That is, tissue that was in the imaging plane during resonance excitation flows out of the imaging plane and is replaced by tissue that was outside of the imaging plane during resonance excitation. The faster the flow, the more complete the replacement of excited tissue with non-excited tissue between excitation and the echo.
Time of flight is based on the physical displacement of a tagged flowing spins with respect to the static spins in the plane of imaging in the time interval between RF excitation and data read-out. That is, resonance is excited in flowing tissue at a selected location to tag it. The location of the tagged tissue a short time later when an echo is generated is noted. The measurements of displacement and time provide an indication of flow.
The flow phase shift technique encodes the flow as an extra complex phase in the final image by means of a flow encoding gradient pulse. See Bryant D. J., Payne J. A., Firmin D. N, and Longmore D. B., J. Comput Assist Tomogr, v. 8, p. 588 (1984); Underwood S. R., Firmin D. N, Klipstein R. H., Rees R. S and Longmore D. B., Br Heart J., v. 57, p. 404 (1987); Klipstein R. H., Firmin D. N, Underwood S. R., Rees R. S and Longmore D. B., Brit. Heart J., v. 58, p. 316 (1987). The flow encoding pulse often consists of an isolated bi-polar or multipolar gradient pulse placed before the read-out acquisition to encode the flow. See Dumoulin, et al., Magn. Reson. Imaging, v. 5, p. 238 (1987). See also FIG. 1 showing a schematic MRI sequence diagram having the bi-polar pulse. The hi-polar pulse can be placed along any one of the three imaging gradients, read, phase and slice.
For phase contrast magnetic resonance imaging in the presence of a magnetic field gradient, moving spins with a velocity value v along the direction of a gradient acquire an extra complex phase over those stationary spins as shown below: ##EQU1## where: .gamma. is the magnetogyric ratio; G(t) is the imaging magnetic gradient field strength of interest at time t measured from the coherent center of the RF excitation pulse coherent center when all the spins are in phase; and .phi..sub.0 is the background phase of stationary spins. Flow acceleration a and pulsatility p are assumed to be small such that their phase contributions can be ignored. This phenomena of phase anomalies in the presence of a magnetic field gradient is used in the gradient waveform design of magnetic resonance imaging sequence (either field echo or spin echo) to measure flow velocity in an in vivo biological system.
For a two-point encoding technique, two different polarities of a bi-polar gradient are used to encode flow velocity in two data acquisitions, which are performed interleaved with respect to each other. The flow raw data of reference and sensitized acquisitions are Fourier transformed and phase reconstructed, respectively. The phase difference between sensitized and reference phase images ideally is proportional to flow velocity (assuming background phase of the phase difference image is zero), i.e.: EQU .DELTA..phi.=.phi..sub.sen -.phi..sub.ref ( 2).
The phase difference image can be directly converted into a flow velocity map V, whose pixel intensity is in cm/sec or other selected units. ##EQU2## where .phi. is the phase difference image in degree, and .DELTA.m.sub.1 is the flow velocity sensitization difference of two sequence pairs in RAD sec/cm. Ideally, without further phase correction on a self-shielded gradient/system, the background (stationary tissue) intensity should be zero (where there is no flow).
The isolated bi-polar flow encoding pulse technique has disadvantages. Placing an isolated bi-polar flow encoding gradient pulse before read-out prolongs the repeat time of the sequence TR and prolongs the echo time TE. The solution of simply integrating the flow encoding gradient with the spatial encoding gradient allows reduction of TE and TR. However, this solution introduces undesirable spatial FOV or slice thickness dependence of the flow encoding. Furthermore, the flow sensitivity of these types of sequences takes only discrete values.
Further, in practice it is difficult to maintain a sufficient dynamical range for a certain flow to be measured. If the flow velocity is outside the effective range of the integrated pulse, the results exhibit the problems of phase wrap and aliasing. To avoid such problems, many flow sequences with different flow sensitivities must be designed to cover a desired range of flow velocities. Each flow sequence would be used over a different velocity range.
The present invention is directed to a new and improved data acquisition technique which overcomes the above-referenced problems and others.